This invention generally relates to ultrasound imaging of the human anatomy for the purpose of medical diagnosis. In particular, the invention relates to methods and apparatus for imaging blood vessel structures, and more particularly, to signal processing algorithms for visualization of blood movement for use in ultrasound imaging systems.
Conventional color flow imaging, including xe2x80x9cangioxe2x80x9d or xe2x80x9cpower Doppler imagingxe2x80x9d (referred to hereinafter as xe2x80x9cflow imagingxe2x80x9d), produces one image from a sequence of transmitted pulses (a packet), typically in the range of 5-15 pulses for each scan line in the image. Slowly moving muscular tissue produces lower Doppler shift in the received signal than signal from moving blood, and efficient clutter filters are designed to suppress the clutter signal to a level much lower than the signal from blood. The signal power after clutter filtering is used to detect points in the image where blood is present. An alternative is to display the signal power as an image (angio or power Doppler) to visualize blood vessels. In order to get reliable detection, substantial temporal and spatial averaging is used, thus limiting the dynamic variation, as well as spatial resolution (bleeding). This averaging process suppresses the spatial speckle pattern in the signal amplitude.
Conventional ultrasound blood flow imaging is based on detection and measurement of the Doppler shift created by moving, scatterers. This Doppler shift is utilized to suppress the signal from slowly moving muscular tissue, in order to detect the presence of blood, and is also used to quantify the actual blood velocity in each point of an ultrasound image. Unfortunately, the Doppler frequency shift is only sensitive to the velocity component along the ultrasonic beam; possible velocity components transverse to the beam are not detected or measurable from the received signal Doppler spectrum. In standard color flow imaging, the Doppler shift is estimated from the received signal generated by a number of transmitted pulses, and coded in a color scale. In some situations, the blood flow direction can be measured from the vessel geometry, but this is difficult to do in an automatic way, especially when the vessel geometry is not clearly visible in the image. Standard color flow imaging often gives confusing blood velocity visualization; e.g., in a curved blood vessel the Doppler shift, and therefore also the color, is changing along the vessel due to change in the angle between the blood velocities and the ultrasonic beam, even though the velocity magnitude is constant. In power Doppler (also called the angio mode) this problem is solved by discarding the measured Doppler shift from the display.
There is considerable interest in measuring the transverse velocity component in ultrasound flow imaging, and a number of methods have been proposed. Compound scanning from two different positions was disclosed by Fox in xe2x80x9cMultiple crossed-beam ultrasound Doppler velocimetry,xe2x80x9d IEEE Trans. Sonics Ultrason., Vol. 25, pp. 281-286, 1978. Compound scanning from two different positions gives two velocity components, but there are practical problems with the large-aperture transducer, the time lag between the measurement of the two components, and the limited field of view. In accordance with a method disclosed by Newhouse et al. in xe2x80x9cUltrasound Doppler probing of flows transverse with respect to beam axis,xe2x80x9d IEEE Trans. Biomed. Eng., Vol. 34, pp. 779-789, October 1987, the transit time through the ultrasound beam is measured, which is reflected in an increased bandwidth of the Doppler signal. This method has very low accuracy, does not yield flow direction, and will only work in regions with rectilinear and laminar flow. Two-dimensional speckle tracking methods based on frame-to-frame correlation analysis have been proposed by Trahey et al. in xe2x80x9cAngle independent ultrasonic detection of blood flow,xe2x80x9d IEEE Trans. Biomed. Eng., Vol. 34, pp. 965-967, December 1987. This method can be used both for the RF signal and the amplitude-detected signal. Coherent processing of two subapertures of the transducer to create lateral oscillations in the received beam pattern has been described by Jensen et al. in xe2x80x9cA new method for estimation of velocity vectors,xe2x80x9d IEEE Trans. Ultrason., Ferroelect., Freq. Contr., Vol. 45, pp. 837-851, May 1998, and by Anderson in xe2x80x9cMulti-dimensional velocity estimation with ultrasound using spatial quadrature,xe2x80x9d IEEE Trans. Ultrason., Ferroelect., Freq. Contr., Vol. 45, pp. 852-861, May 1998. This method gives quantitative lateral velocity information, including the sign. The main drawback of this method is poor lateral resolution, which limits its use for imaging.
There is a need for a method of ultrasound imaging which gives the system user a correct perception of the blood flow direction and magnitude, and which is also useful to separate true blood flow from wall motion artifacts.
In ultrasound imaging, the returned echoes are processed coherently. In the images there are variations in the intensity due to constructive and destructive interference of the sound waves scattered back from a large number of scatterers. These variations in the intensity is often termed the xe2x80x9cspeckle patternxe2x80x9d. When there is a slight displacement of the scatterers (red blood cells), there will be a corresponding displacement of the speckle pattern. By enhancing the speckle pattern from moving scatterers and display a stream of such images, an intuitive display of the blood flow is obtained.
The present invention comprises a method and an apparatus for imaging blood motion by preserving, enhancing and visualizing speckle pattern movement, which is related to the blood cell movement in the blood vessels. This method will be referred to herein as xe2x80x9cblood motion imagingxe2x80x9d (BMI). Speckle pattern movement gives the user a correct perception of the blood flow direction and magnitude, and is also useful to separate true blood flow from wall motion artifacts. In this way, the system operator can see the blood flowing in the image, although no attempt is made to measure the lateral velocity component. However, the lateral velocity component may be derived indirectly by combining an angle measurement derived from the speckle motion with the radial velocity component obtained from the Doppler frequency shift.
In one preferred embodiment of the invention, multiple image frames per packet of transmitted pulses are produced, instead of a single image frame. The motion of the blood scatterers creates a corresponding movement of the speckle pattern in the images from frame to frame, showing both radial and lateral movement. The time between each of these frames equals the pulse repetition time (1/PRF) within the signal packets. In order to visualize the motion, the display frame rate must be reduced substantially, e.g., from 1 kHz to 30 Hz. For real-time display, much data must be discarded, but for slow motion replay, a larger fraction or all of the recorded frames can be used.
In accordance with the preferred embodiment of the invention, the data are acquired as in conventional color flow imaging. A series of pulses (a packet) are transmitted in each beam direction and echoes are acquired for a region of interest (ROI) in the blood motion image. The pulse firings within a packet are separated by a constant time interval. This time interval is much smaller than the time between successive packets. Then one tissue image, which may extend beyond the blood motion image ROI, is recorded. The maximum possible pulse repetition frequency (PRF) during packet acquisition is determined by the imaging depth. By reducing the PRF, it is possible to use a technique called beam interleaving. After firing a pulse in a first direction, there is time available to fire pulses in one or more different directions before firing the next pulse in the first direction. This collection of beam directions is called an interleave group. By using a relatively broad transmit beam, it is possible to acquire several receive beams per transmit beam by simultaneous beamforming in slightly different directions. This known technique is called multi-line acquisition (MLA).
The data input for signal processing are the beamformed and complex-demodulated I/Q data samples. Alternatively, the processing can be performed on the real-valued RF data without complex demodulation. In accordance with the processing technique disclosed herein, several images per packet are displayed, as opposed to conventional color flow imaging in which only one image per packet is displayed. The first step in the BMI processing is high-pass filtering of the signal vector from each range gate. Following high-pass filtering, the speckle signal is formed. The speckle signal is then subjected to a nonlinear scale conversion. An example of this is logarithmic compression followed by gain and dynamic range adjustment. The resulting speckle signal is displayed as the desired blood motion image concurrently with a corresponding tissue image.
In accordance with another preferred embodiment, fluctuation in the mean power from packet to packet is compensated for in order to obtain a smooth temporal display. This is accomplished by dividing each speckle signal sample by the mean value calculated for the packet, thereby forming an enhanced speckle signal for imaging blood motion. In the log domain this is equivalent to subtracting the logarithm of the mean value from the logarithm of each speckle signal sample.
In general, time averaging will reduce the speckle variation. However, time averaging within one packet will produce a trace pattern in the image along the blood flow direction which show the direction of flow even in a still frame image. Further time averaging (between packets), which is usually done in conventional color flow imaging, will destroy this trace pattern.
Therefore, in accordance with a further preferred embodiment, the moving speckle patterns are processed temporally. As a result of the temporal processing, the moving speckle patterns create traces in the image along the streamlines in the flow. Simple temporal averaging within each packet gives a non-directive streamline effect. More sophisticated methods can also preserve direction information, and to some extent velocity magnitude. Temporal processing is not necessary for the visual perception of flow, but makes possible still frame visualization of flow direction and magnitude, and may give improvement for real-time display, where the frame rate must be limited. If the smoothing window for temporal averaging is chosen equal to the packet size, and one image is generated for each packet, the processing would be similar to standard color flow imaging. However, a number of steps may preferably be taken to accentuate the speckle pattern in the flow image. First, the spatial resolution should be as high as possible by using a short transmitted pulse and a large-aperture transducer. Second, the number of scan lines per beam interleave group should be as high as possible. This can be achieved by using a low PRF and/or MLA (parallel receive beams). Third, the speckle signal can be normalized by a local average obtained by temporal and/or spatial averaging.
There are various ways of including the speckle pattern in the flow image: (1) by combining the speckle signal with the signal power and showing the combined signal in the same way as the angio mode flow image; (2) by intensity (xe2x80x9cvaluexe2x80x9d in HSV color representation) modulation of the color flow or angio image; (3) by color coding the xe2x80x9cagexe2x80x9d of the speckle, in order to visualize the direction and the magnitude of the movement.
The invention can be implemented as post-processing, based on recorded I/Q data of a sequence of images, or in real-time. The invention can be implemented in hardware or software.